The present disclosure is related to the field of stents, and more specifically to an implantable, sutureless bioprosthetic stent graft comprising a biomaterial. The disclosure is further related to a device and a method for suturelessly bonding a biomaterial to a bioprosthetic frame of a stent graft.
Development of tissue substitutes has been undertaken for replacement and repair of damaged or diseased tissue and organs. Where there is a lack of native tissue, reconstruction generally is performed with an autograft, heterogaft and allograft. Various tissue grafts and synthetic biomaterials typically have been unsuccessful based on mechanical, structure, functional, or biocompatibility problems. There is therefore no ideal biomaterial for tissue replacement, particularly for soft tissue and tubular organs (e.g., vascular, trachea, esophagus, and biliary tract, and urinary tracts tissue).
Prosthetic stents and valves have been described in the prior art. Stents have been used with some success to overcome the problems of restenosis or re-narrowing of a vessel wall. Stents are exemplified by U.S. Pat. No. 6,293,968 (to Taheri) and U.S. Pat. No. 5,306,286 (to Stack et al.), which teaches a prosthetic stent constructed of synthetic materials. U.S. Pat. No 6,293,968 and U.S. Pat. No 5,306,286.
However, the use of such devices is often associated with thrombosis and other complications. Additionally, prosthetic devices implanted in vascular vessels can exacerbate underlying atherosclerosis.
Medical research therefore has focused on trying to incorporate artificial materials or biocompatible materials as bioprosthesis coverings to reduce the untoward effects of metallic device implantation, such as intimal hyperplasia, thrombosis and lack of native tissue incorporation.
Synthetic materials for stent coverings vary widely, e.g., materials such as Gore-Tex®, polytetrafluoroethylene (PTFE), and a resorbable yarn fabric (U.S. Pat. No. 5,697,969 to Schmitt et al.). Synthetic materials generally are not preferred substrates for cell growth.
Biomaterials and biocompatible materials also have been utilized in prostheses. Such attempts include a collagen-coated stent, taught in U.S. Pat. No. 6,187,039 (to Hiles et al.). As well, elastin has been identified as a candidate biomaterial for covering a stent (U.S. Pat. No. 5,990,379 (to Gregory)).
In contrast to synthetic materials, collagen-rich biomaterials are believed to enhance cell repopulation and therefore reduce the negative in vivo effects of metallic stents. It is believed that small intestinal submucosa (SIS) is particularly effective in this regard.
Some of the above-discussed coverings, while used to prevent untoward effects, actually exacerbate the effects to some extent. Accordingly, it is desirable to employ a native biomaterial or a biocompatible material to reduce post-procedural complications.
Mechanically hardier stent graft devices are required in certain implantation sites, such as cardiovascular, aortic, or other locations. In order to produce a sturdier bioprosthetic stent, a plurality of layers of biomaterial typically are used. Suturing is a poor technique for joining multiple layers of biomaterial. While suturing is adequate to join the biomaterial sheets to the metallic frame, the frame-sutured multiple sheets are not joined on their major surfaces and are therefore subject to leakage between the layers. Suturing of the major surfaces of the biomaterial layers also introduces holes into the major surfaces, increasing the risk of conduit fluid leaking through or a tear forming in one of the surfaces.
Heretofore, biomaterials have been attached to bioprosthetic frames using conventional suturing techniques. However, this approach is disadvantageous from manufacturing and implantation perspectives.
Suturing is time-consuming and labor-intensive. For example, suturing a sheet of biomaterial over a stent frame typically is a one- to two-hour process for a trained person and of the covered stents made, many are rejected. It is also an operator dependent process that can lead to issues with product uniformity and reliability. As well, suturing entails repeatedly piercing the biomaterial, creating numerous tiny punctures that can weaken the biomaterial and potentially lead to leakage and infection after the graft device has been installed.
Moreover, the presence of suture material can enhance the foreign body response and lead to tubular vessel narrowing at the implantation site.
As an alternative to suturing, U.S. Pat. Nos. 5,147,514, 5,332,475, and U.S. Pat. No 5,854,397 describe processes for photo-oxidizing collageneous material in the presence of a photo-catalyst to crosslink and stabilize the collageneous material. Reconstituted soluble collagen fibrils are taught to be mixed and suspended in solutions containing a photo-catalyst, so that a photo-oxidizative cross-linking process can be performed to produce stabilized collagen products.
However, the references fail to teach crosslinking of collagen fibrils between two individual native tissues, as well as fusion of those separate tissue pieces using photo-oxidization techniques.
Biocompatible adhesive compounds also have been investigated as alternatives to suturing. For example, fibrin glue, a fibrinogen polymer polymerized with thrombin, has been used as a tissue sealant and hemostatic agent.
Bioadhesives generally produce rigid, inflexible bond regions that can lead to local biomaterial tears and failure of the graft device. In addition, some bioadhesives and photochemical cross-linking agents (e.g., glutaraldehyde) carry risk of acute and chronic toxicity and biocompatibility.
Bio-tissue welding, using a laser, is known in the art, e.g., U.S. Pat. No. 5,156,613 (to Sawyer). This technique uses light energy to heat an area of tissue sufficiently to denature at least a portion of the tissue constituents and fuse them together.